Stents are endovascular prostheses which may be used for the treatment of stenoses (vasoconstriction). Stents have a tubular or hollow cylindrical basic mesh which is open at both longitudinal ends. The tubular basic mesh, composed of the base material of such an endoprosthesis, is inserted into the blood vessel to be treated and is used to support the vessel.
Such stents have become established for the treatment of vascular diseases in particular. Use of stents allows constricted regions in blood vessels to be expanded, resulting in lumen gain. Although the optimal vessel cross section primarily necessary for successful treatment may be achieved by the use of stents, the permanent presence of such a foreign body initiates a cascade of microbiological processes which may lead to gradual overgrowth of the stent, and in the worst case may result in vascular occlusion. A starting point for solving this problem consists in producing the stent from a biodegradable material.
The term “biodegradation” refers to hydrolytic, enzymatic, and other metabolic chemical degradation processes in the living organism which are primarily caused by the bodily fluids which come into contact with the endoprosthesis, resulting in gradual dissolution of at least large portions of the endoprosthesis. The term “biocorrosion” is often used synonymously for “biodegradation.” The term “bioabsorption” includes the subsequent absorption of the degradation products by the living organism.
Suitable materials for the basic mesh of biodegradable endoprostheses may be of a polymeric or metallic nature, for example. The basic mesh may also be composed of several materials. These materials share the common feature of biodegradability. Examples of suitable polymeric compounds include polymers from the group comprising cellulose, collagen, albumin, casein, polysaccharides (PSAC), polylactide (PLA), poly-L-lactide (PLLA), polyglycol (PGA), poly-D,L-lactide-co-glycolide (PDLLA-PGA), polyhydroxybutyric acid (PHB), polyhydroxyvaleric acid (PHV), polyalkyl carbonates, polyortho esters, polyethylene terephtalate (PET), polymalonic acid (PML), polyanhydrides, polyphosphazenes, polyamino acids, and the copolymers thereof, as well as hyaluronic acid. Depending on the desired characteristics, the polymers may be present in pure form, derivatized form, in the form of blends, or as copolymers. Metallic biodegradable materials are based on alloys of magnesium, iron, zinc, and/or tungsten. The present invention preferably relates to stents or other endoprostheses whose biodegradable material contains magnesium or a magnesium alloy, particularly preferably the alloy WE43, and/or a biodegradable polymer, particularly preferably PLLA. As base materials for stents, these materials have a particularly suitable combination of mechanical, biological, and corrosive properties.
Stents which have coatings for various functions are presently known. Implementation of biodegradable implants involves the problem of controlling the degradability corresponding to the intended therapy. It has not been possible thus far to design a stent which loses its integrity within the target corridor of four weeks to six months, which is important for many therapeutic applications. In this regard “integrity,” i.e., mechanical integrity, refers to the characteristic that the stent or endoprosthesis does not undergo hardly any mechanical losses compared to the undegraded stent. This means that the stent is still mechanically stable enough to ensure that the collapse pressure drops only slightly, i.e., to a maximum of 80% of the nominal value. Thus, when integrity is present the stent is still able to meet its primary function of keeping the blood vessel open. Alternatively, integrity may be defined such that the stent is mechanically stable enough that in a load state in the blood vessel it undergoes minimal changes in its geometry, for example does not show appreciable collapse, i.e., under a load of at least 80% of the dilation diameter, or has very little breakage of supporting struts.
Degradable magnesium stents have proven to be particularly promising for the referenced target corridor of degradation, although on the one hand they lose their mechanical integrity or supporting effect too soon, and on the other hand show great fluctuations in loss of integrity both in vitro and in vivo. This means that for magnesium stents the collapse pressure drops too rapidly over time, and/or the drop in the collapse pressure varies too greatly and therefore cannot be determined.
Basically, there are three known approaches for adjusting the desired target time window for the loss of integrity. First, a thicker optimized stent design may be selected. Second, an optimized, slowly degrading magnesium alloy may be used for the stent, or third, surface layers may be provided which delay or accelerate the onset of degradation of the magnesium basic mesh, and/or influence the point in time that degradation begins. The possibility of varying the degradation characteristics according to the first or second approach is greatly limited, and may not be sufficient for an economical and clinically satisfactory solution. With regard to the first approach, in order to ensure ease of insertion of the stent and due to the limited blood vessel dimensions, wall thicknesses of greater than 200 μm are not advisable. For the second approach, only a very limited spectrum of biocompatible and moderately rapidly degradable alloys is known. With regard to the third approach, only fluorine passivation is known.
The above-referenced passivation layers have two fundamental disadvantages which result, among other reasons, from the fact that such stents usually assume two states, namely, a compressed state with a small diameter and an expanded state with a larger diameter. In the compressed state the stent can be inserted into the blood vessel to be supported by using a catheter, and positioned at the site to be treated. The stent is then dilated at the treatment site using a balloon catheter, for example, or, if a shape memory alloy is used as stent material, is converted to the expanded state, for example by heating above the transition temperature. As a result of this change in diameter the basic mesh of the stent is subjected to mechanical stress. Additional mechanical stresses on the stent may occur during manufacturing, or when the stent moves in or with the blood vessel in which the stent is inserted. Thus, the referenced passivation results in the disadvantage that during deformation of the implant microcracks are produced which lead to infiltration of the coating material, thereby reducing the passivation effect of the coating, which causes nonspecific localized degradation. In addition, the onset and speed of degradation depend on the size and distribution of the microcracks, which are defects that are difficult to monitor. This results in a large dispersion in the degradation times.
WO 2005/065576 A1 discloses control of the degradation of degradable implants by use of a coating made of a biodegradable material. Location-dependent degradation of the implant is optimized by the fact that the base body has an in vivo location-dependent first degradation characteristic and a coating which covers the base body completely or only in places and consists of at least one biodegradable material, the coating having an in vivo second degradation characteristic. The cumulative degradation characteristic at a location is obtained from the sum of the respective degradation characteristics of the material and the coating at the given location. The location-dependent cumulative degradation characteristic is specified by varying the second degradation characteristic in such a way that the degradation takes place at the given location during a predetermined time period at a predeterminable degradation rate.
The degradation characteristic of the biodegradable coating described in WO 2005/065576 A1 is achieved in a very general manner, in particular by varying the morphological structure of the coating, by substantive modification of the material, and/or by adjusting the layer thickness of the coating. In this regard “morphological structure” is understood to mean the conformation and aggregation of the compounds which form the coating. The cited document references hyaluronic acid as an example of a coating.
U.S. 2006/0224237 A1 likewise describes a transplant or stent having a protective layer which is used to protect surface structures of the stent from destruction. The surface structures may be formed from one or more materials which are at least partially dissolved, degraded, or absorbed under various environmental conditions.
The possibilities stated in the cited documents for influencing the degradation do not include satisfactory approaches for endoprostheses which degrade within the referenced target corridor. WO 2005/065576 A1 describes only very general principles which do not provide specific approaches in particular for magnesium stents.
U.S. 2007/0050009 A1 concerns a stent having a support structure composed of biodegradable material. This support structure is at least partially provided with an absorption inhibitor layer which reduces the rate of absorption of the support structure. The absorption inhibitor layer itself is likewise absorbed by the surrounding bodily fluids. This known approach as well provides only very limited control of degradation of the stent, which for many applications is inadequate. Hyaluronic acid, collagen, or polyglycolic acid are referenced as examples of materials for an absorption inhibitor layer.
DE 10 2005 039 126 A1 and U.S. 2005/0196424 A1, among other sources, describe coatings with parylene as a protective layer, in particular for prevention of restenosis or inflammation after implantation or as a pretreatment layer for a carrier of bioactive materials. Use of these coatings for the control of degradation is not disclosed.